Typically, in computed tomography (CT) imaging systems, a rotatable gantry includes an x-ray tube, detector, data acquisition system (DAS), and other components that rotate about a patient that is positioned at the approximate rotational center of the gantry. X-rays emit from the x-ray tube, are attenuated by the patient, and are received at the detector. The detector typically includes a photodiode-scintillator array of pixelated elements that convert the attenuated x-rays into photons within the scintillator, and then to electrical signals within the photodiode. The electrical signals are digitized and then received within the DAS, processed, and the processed signals are transmitted via a slipring (from the rotational side to the stationary side) to a computer or data processor for image reconstruction, where an image is formed.
The gantry typically includes a pre-patient collimator that defines or shapes the x-ray beam emitted from the x-ray tube. X-rays passing through the patient can cause x-ray scatter to occur, which can cause image artifacts. Thus, x-ray detectors typically include an anti-scatter grid (ASG) for collimating x-rays received at the detector. Imaging data may be obtained using x-rays that are generated at a single polychromatic energy. However, some systems may obtain multi-energy images that provide additional information for generating images.
Third generation multi-slices CT scanners typically include a detector assembly having scintillator/photodiodes arrays positioned in an arc, where the focal spot is the center of the corresponding circle. The material used in these detectors generally has scintillation crystal/photodiode arrays, where the scintillation crystal absorbs x-rays and converts the absorbed energy into visible light. A photodiode is used to convert the light to an electric current. The reading is typically proportional and linear to the total energy absorbed in the scintillator.
In X-ray computed tomography (CT) imaging systems, the x-ray tube generates high speed electrons from the filament. The electrons fly toward the positive target anode, in which the energy of the electrons is converted to X-rays. In conventional CT scanners, the X-ray emits from one focal spot on the anode plate. For multi-row scanners, to increase the resolution and reduce or remove under-sampling related image artifacts the so-called “Flying focal spot” (FFS), i.e. the focal spot is periodically moved among certain given positions, can be employed. The in-plane focal spot motion can increase resolution of transverse planes while the motion in the z-direction, referred to hereinafter as zFFS, which can increase axial resolution.
In a conventional single focal spot cone beam (CB) system, the sampling interval in the z-direction is the same as detector height of each row. The practical detector height is constrained by production technology and cost, and can cause windmill artifacts of helical scans in high contrast regions. zFFS strategy can increase the sampling rate in the z-direction, thus it not only can boost the z-resolution but also reduce helical windmill artifacts. According to this disclosure a fundamental circular cone beam (CCB) scan protocol is used to present the disclosed method, but it is contemplated that the disclosure is applicable to helical scans, as well.
X-ray Tomography is widely used in clinical disease diagnosis. The zFFS strategy has been proposed for several years and CT scanner venders have produced products to implement the focal spot wobbling idea. In general, current image reconstruction methods for zFFS scanning treat detector readings from alternating two focal spots as interleaved sampling, i.e. group the two sets of data to one set by interleaving the rows of each consecutive (odd and even) reading pair to build one sinogram with double number of rows. Then, the combined data is used for image reconstruction by a regular single focal spot geometry, either using a native fan geometry, or by re-binning the data to a parallel geometry. This type of reconstruction method, using combined data, has at least two drawbacks:
1) Small FOV: rebuilding data sets by interleaving assumes that the rays from two focal spots are stacked alternatively in a z-direction, which is true for a limited FOV. This assumption does not hold for the voxels close to focal spots. And, in fact, in one example the assumption only holds for a field-of-view (FOV) of about 200 mm for some known commercial scanners.
2) Inaccuracy: For voxels out of the limited FOV the interleaved data may cause mistakes. And, for voxels within the limited FOV the interleave strategy may also introduce inaccuracy since a geometrically “perfect” and equally spaced z-interleave for the combined data only happens at the z-axis. That is, the further the voxel from the z-axis, the worse the violation of the equal space assumption.
Thus, there is a need to improve zFFS reconstruction algorithms for both CCB and helical scans.